Nuclear spin imaging is a new non-destructive investigation method, one of the most important fields of application thereof being medical diagnostics. The principle of nuclear spin imaging was presented by P. Lauterbur in 1973. (Lauterbur: Nature vol. 242, Mar. 16, 1973, p. 190 . . . 191). Prior to this, the operating principle of a kind of NMR phenomenon-based investigation apparatus was proposed by R. Damadian. (Damadian: U.S. Pat. No. 3,789,832). A plurality of nuclear spin imaging methods have been developed and described e.g. in references Ernst: U.S. Pat. No. 4,070,611, Garroway et al: U.S. Pat. No. 4,021,726 and Moore et al: U.S. Pat. No. 4,015,196.
Nuclear spin imaging, as well as other NMR investigation methods, are based on the fact that nuclei of some elements have a magnetic moment. Such elements include e.g. hydrogen, fluorine, carbon and phosphorus with certain isotopes thereof having a nuclear magnetic moment. Let us study e.g. the nucleus of a hydrogen atom, i.e. the proton, which is a positively-charged primary particle. Proton rotates around its own axis, i.e. it has a certain spin. The rotation creates the magnetic moment of a proton and also the flywheel moment parallel to the axis of rotation.
If a number of hydrogen atoms are placed in an external magnetic field B.sub.o, most of the magnetic moments of said nuclei settle parallel to the external field B.sub.o and thus there is developed in the bunch of hydrogen atoms a net magnetization M.sub.n, which is directly proportional to said external magnetic field B.sub.o. However, the temperature of the subject bunch of atoms has an effect on how large is the majority creating the net magnetization as compared to the entire bunch of nuclei. When the temperature of the object is e.g. that of a human body, the quantitative difference between the majority and minority in a bunch of nuclei is just 1 millionth of the total number of nuclei. If temperature of the object could be lowered, net magnetization would increase inversely proportionally to the absolute temperature of said object.
In pulse NMR investigations, the resulting net magnetization M.sub.n is deflected with a strong radio-frequency magnetic pulse 90.degree. from the direction of said external magnetic field B.sub.o. As a result of the interaction between a flywheel moment as well as a magnetic moment, created by spring of the nuclei, and an external field, the resulting net magnetization is set in precession. The angular speed of precessive magnetic moment is directly proportional to the external magnetic field according to formula 1 EQU W.sub.o =.gamma.B.sub.o ( 1)
wherein
.gamma. is gyromagnetic ratio PA1 B.sub.o is strength of external magnetic field PA1 W.sub.o is so-called Larmor frequency PA1 k is a coefficient independent of field PA1 N is speed of rotation of detection coil PA1 A is cross-sectional area of the coil PA1 f filling ratio PA1 Q is quality factor of the coil PA1 W.sub.o is Larmor frequency PA1 L is inductance of the coil PA1 B is the bandwidth employed
If outside an object or a target is placed an induction coil and a capacitor for providing a resonance circuit, the magnetization in precession will induce a signal voltage in the terminals of said resonance circuit. The amplitude of signal voltage V.sub.s is directly proportional to the Q-factor or quality factor of a resonance circuit.
A more important quantity than signal voltage is the signal/noise ratio SNR. Nuclear spin imaging like all other NMR investigations depend on the obtainable signal/noise ratio. If the electric losses of a target to be examined are ignored, the resulting signal/noise ratio will be: EQU SNR=kNAf(QW.sup.3.sub.o /LB).sup.1/2 ( 2)
wherein
As set forth in equation 2, the signal/noise ratio achieved in NMR imaging is inversely proportional to the square root of a bandwidth. In nuclear spin imaging a target is covered by a magnetic field gradient during signal collection. If the order of a gradient is g [T/m] and the projection of a target in the direction of gradient is 1[m] in length, the frequency bandwidth BW of an NMR signal inducing a target in the signal coil is EQU BW=(.gamma./2.pi.).multidot.g.multidot.1 (3)
Typically 1=0.2 m if the target to be imaged is the head of a human being and 1=0.5 m if the target is the thorax region of a human body. The gradients used are generally in the order of g=1 mT/m. Thus, the bandwidth of a signal obtained from a head size target is BW.apprxeq.8 kHz and from the thorax region BW.apprxeq.20 kHz. In normal hospital conditions, the same apparatus is required to successively take images of targets of varying sizes: body imaging being performed immediately after head imaging. If the strength of gradients used in both imagings is constant, the bandwidth of a signal in body imaging will be approximately two-fold and the necessary signal catch band will also be two-fold. Hence, when proceeding to head imaging, it is preferable to reduce the bandwidth of signal catch or collection for the improvement of signal/noise ratio.
According to the prior art, it is conventional to employ different frequency bandwidths with the assembly tuned for head imagings or body imagings. The bandwidths applied in both imagings must thus be dimensioned in a manner that all relevant targets to be imaged can be imaged. Thus, the bandwidth used in most cases is too large and therefore the signal/noise ratio will be lower than what it could be at its optimum.